1. Field of the Invention
The present invention relates to a method for compensating x-ray imaging systems for radiation scatter.
2. Description of the Prior Art
In x-ray imaging systems, the total flux of detected radiation consists not only of photons that did not interact with the elements of the attenuating object being imaged but also of radiation scatter. Specifically for systems using area detectors the amount of radiation scatter can be very large. For a large class of applications, such as energy-selective imaging, this radiation can be an important source of error and must be compensated in order to have satisfying results.
A lot of work has been done already in searching good methods for compensating x-ray imaging systems for radiation scatter.
The use of grids or air gaps reduces but doesn't eliminate the radiation scatter. For many applications, as for dual-energy subtraction imaging, it is not sufficient.
Several analytical models, representing the scatter--some of them use point spread functions--are proposed in the prior art. They require parameters the values of which are difficult to obtain and in most cases to be found experimentally. These models do not give satisfying results in some practical applications.
In the prior art, some investigators (Molloi SY, Mistretta Calif. Scatter-glare correction in quantitative dual-energy fluoroscopy. Med. Phys. 1988; 15:289-297) use correction tables for a specific application that give a hypothesized relationship between the detected grey level in a certain pixel and the scatter fraction. For most applications they are rough estimations with insufficient accuracy.
In U.S. Pat. No. 4,549,307 a method is used in which two x-ray irradiations of an object are made: one with a disk sampler, consisting of an array of small lead beam stops above the object, and one without the sampler. In the shadow of each disk only scattered radiation is detected and the average of the pixels values in this shadow gives the value of the radiation scatter at that location. By interpolation of the sample values of the radiation scatter distributed over the image area, an estimation of the radiation scatter values in the whole image is generated. The scatter corrected image is obtained by subtracting the scatter-surface from the second image, where no beam stop array was positioned. One disadvantage of this method is that one needs two separate images, and thus two separate shots, of the object. In medical applications, this means that the patient (=the object) receives a larger x-ray dose and that he may have moved between the two shot's. Switching quickly detectors and disk samplers asks for a mechanically complicated system. Another implementation of the method with the lead beam stops in which only the x-ray irradiation with the disk sampler is made has the disadvantage that all the information about the object is lost under the beam stops. This can be an important disadvantage.
In the prior art, another method for scatter radiation compensation (Shaw C. A novel technique for simultaneous acquisition of primary and scatter image signals. SPIE Vol. 1651 Medical Imaging VI: Instrumentation (1992), p. 222-233), is the so-called primary-modulation-demodulation technique. The primary x-ray distribution is modulated and demodulated with two filters of equal material and thickness placed on the tube and detector sides of the objects. The modulation-demodulation process results in a reduction of scatter signals in selected regions of the image. It leaves the overall primary distribution signal unchanged. The signal drop of the radiation scatter can be measured and used to estimate the scatter radiation signal in the selected regions. Although the PMD method allows both primary and scatter signals to be acquired simultaneously, it has two main disadvantages: it is unknown how the drop in scatter radiation relates to location, scattering geometry, object (mostly a patient), etc.. Another disadvantage is that it is practically impossible to match the modulator and the demodulators. Therefore the results are based on rough estimates and the accuracy is reduced.
A recently developed technique is described by C. Fivez in unpublished European application no. 93.203671. The methods according to that invention are based on one irradiating shot of the object. After scatter radiation compensation, the information about the object is not lost in any location of the object image. By comparing the detected signal under a partially transparent body (e.g. disk or strip), positioned between the x-ray source and the object being imaged, with the signal in the image near the border of the shadow of the partially transparent body, the radiation scatter signal in the location of the shadow of the body is calculated. In case of a polychromatic source, calibration with two known materials allows accurate calculation of the radiation scatter. The partially transparent bodies are positioned at several locations between object and source and, by means of an interpolation technique, the radiation scatter in every location of interest can be calculated. The radiation scatter image is subtracted from the original image of the object. The primary signal (=without radiation scatter) in the location of the bodies has undergone an extra drop because of the partially transparent body but the information about the object under the body is not lost. The method allows, with only one shot of the object, to compensate for radiation scatter without loosing information about the object. For accurate calculation of the scattered radiation one needs a calibration as mentioned before. In energy-selective imaging, one often uses such calibration, so that in such cases it is not a problem. Nevertheless, in other application areas, the calibration can be an objection, for it complicates the procedure. Another disadvantage of the method is that the materials in the object and of the partially transparent bodies must belong to (a large) group of materials. It can not be whatever materials. For the partially transparent bodies a consequence can be that the height of the bodies must be relatively large, if one wants a reasonable accuracy. Such heights can cause some geometric artefacts.
U.S. Pat. No. 4,688,242 discloses a X-ray imaging system wherein X-ray image data are calculated from scattered X-ray intensity distribution data and transmission X-ray data.
The transmission X-ray image data are obtained by irradiating an object with X-rays and detecting the transmitted rays by means of an X-ray image detection means such as an image intensifier and a camera.
The scattered X-ray intensity distribution data are calculated on the basis of (i) a plurality of transmission X-ray data obtained by irradiating an object with X-rays with a mask member, comprising a plurality of X-ray shielding regions distributed in a predetermined pattern being located at different positions in said X-ray radiation field and (ii) on transmission X-ray data obtained by irradiating said object with X-rays when said mask member is located outside said X-ray radiation field.
The method applied in this X-ray imaging system comprises (1) the insertion of the mask member between the X-ray image detection means and the X-ray source and (2) the shifting of the mask member between (two) predetermined positions in the X-ray irradiation field. In particular the movement of the mask means from a position outside the X-ray radiation field to a position in between the X-ray source and the object is time consuming and inadequate for application in chest imaging since the position of the patient could have changed within the period of time required to perform the displacement of the shield. The method is further not adapted for application in a method wherein an area detector such as a photostimulable phosphor screen is used.
In the article by F. Wagner, A. Macovski and D. Nishimura entitled `Dual Energy X-ray projection imaging: Two sampling schemes for the correction of scattered radiation`, published in Medical Physics, Vol. 15, Oct. 1988, p. 732-748 two methods have been disclosed for the correction of scatter.
Each method requires two measurements and each method involves placing an opaque sampling grid between the source and the object.
In the first method, the grid is an array of lead disks present during one measurement. Using this grid, an estimate of the scatter field is generated and subtracted from the second measurement yielding a scatter-corrected image. This method also comprises the time-consuming step of positioning the grid in between source and object, so it is also inadequate for most chest applications.
In the second described method the grid is present during both measurements and is shifted by one-half of a strip spacing to completely sample the image.
An image to be corrected is generated, an estimate of the scatter field and a scatter corrected image are generated. This method is disadvantageous because at locations beneath a lead disk the original signal is lost.